This application relates to methods for correction of high-order primary decay in detectors, and more particularly to methods providing increased flexibility in defining shapes of corrected decay curves in detectors.
As used herein, the term xe2x80x9cprimary decayxe2x80x9d refers to the fastest exponential decay component of a scintillator. xe2x80x9cAfterglowxe2x80x9d refers to remaining, slower decay components.
In at least one known computed tomography (CT) imaging system configuration, an x-ray source projects a fan-shaped beam which is collimated to lie within an X-Y plane of a Cartesian coordinate system and generally referred to as the xe2x80x9cimaging planexe2x80x9d. The x-ray beam passes through the object being imaged, such as a patient. The beam, after being attenuated by the object, impinges upon an array of radiation detectors. The intensity of the attenuated beam radiation received at the detector array is dependent upon the attenuation of the x-ray beam by the object. Each detector element of the array produces a separate electrical signal that is a measurement of the beam attenuation at the detector location. The attenuation measurements from all the detectors are acquired separately to produce a transmission profile.
In known third generation CT systems, the x-ray source and the detector array are rotated with a gantry within the imaging plane and around the object to be imaged so that the angle at which the x-ray beam intersects the object constantly changes. A group of x-ray attenuation measurements, i.e., projection data, from the detector array at one gantry angle is referred to as a xe2x80x9cviewxe2x80x9d. A xe2x80x9cscanxe2x80x9d of the object comprises a set of views made at different gantry angles, or view angles, during one revolution of the x-ray source and detector. In an axial scan, the projection data is processed to construct an image that corresponds to a two dimensional slice taken through the object. One method for reconstructing an image from a set of projection data is referred to in the art as the filtered back projection technique. This process converts the attenuation measurements from a scan into integers called xe2x80x9cCT numbersxe2x80x9d or xe2x80x9cHounsfield unitsxe2x80x9d, which are used to control the brightness of a corresponding pixel on a cathode ray tube display.
It is known to use europium-containing polycrystalline ceramic scintillators in detector arrays of computed tomographic (CT) imaging systems. Such scintillators exhibit much lower hysteresis and radiation damage than other known solid state detectors, such as CdWO4. In addition, the detector material is highly transparent, resulting in higher light output. Detection quantum efficiency (DQE) of 98% or more at 3 mm depth is obtained in a clinical x-ray energy range, resulting in improved image quality. Nevertheless, output signal decay, as indicated by its primary decay, is relatively slow (close to 1 millisecond), discouraging use of these otherwise advantageous scintillators at fast sampling rates at high scanning speeds. An exemplary decay curve is shown in FIG. 3.
It has been shown that a slow primary speed of the detector degrades spatial resolution of a CT imaging system, especially at higher scanning speeds. For example, a scan at 0.5 seconds per rotation will be degraded relative to a scan at 1.0 seconds per rotation, resulting from the significantly increased sampling rate. To overcome this shortcoming, recursive correction algorithms have been proposed.
Corrections using recursive correction algorithms compensate not only for the effects of the primary speed component of the detector response, but also for the afterglow components. Known techniques perform satisfactorily for scan speeds up to 1.0 second using scintillation materials having a decay characteristic such as that illustrated in FIG. 3. For faster scan speeds. however, undershoot and overshoot on the decay curve will occur, resulting in streak artifacts in reconstructed images. This undershoot and overshoot phenomenon is not an intrinsic feature of the decay of this scintillation material, but rather is a side effect of the signal restoration technique.
It would therefore be advantageous to provide methods to reduce or void undershoot and overshoot of the decay curve and the resulting artifacts in the reconstructed images.
There is therefore provided, in one embodiment, a method for calibrating a primary decay correction for a radiation detector having a decay curve that can be characterized by a plurality of components having different time constants. The method includes steps of fitting the decay curve to a sum of a plurality of weighted exponentials having a first set of time constants; applying a correction to a measured response of the detector using a sum of the plurality of weighted exponentials having the first set of time constants to obtain a corrected response; selecting at least one additional exponential time constant dependent upon the corrected response; and fitting the decay curve to a sum of a second plurality of weighted exponentials including the first plurality of time constants and the at least one additional exponential time constant.
The above described embodiment reduces or avoids undershoot and overshoot of the decay curve and the resulting artifacts in reconstructed images.